Valve replacements are among the most widely used cardiovascular devices and the demand for them is increasing. At the present time the clinical devices available are limited to mechanical and biological valves. Long-term clinical applications of such valves are however highly problematic, given some persistent critical problems such as thrombogenicity and service life.
In fact mechanical valves have a service life and do not need repeat surgery, in that they are not subject to structural failures, but because they give rise to thromboembolic complications patients have to take anticoagulant treatments for their entire lives. Biological valve prostheses made of porcine, bovine or equine pericardium modelled and sutured onto support structures (stents) reproduce the functional biomechanical characteristics of native valves, give rise to fewer thromboembolic complications but in many cases have to be replaced 10-15 years after implant because of the occurrence of calcification problems and damage brought about by the decellularisation treatments undergone in order to reduce problems associated with immunological response.
The use of xenografts (whole valves taken from animals) and homografts (whole valves taken from cadavers) has hitherto been limited by rejection problems and very low availability.
Polymeric heart valves (hereinafter PHV) have been investigated for a long time, but their success has been impeded by very short service life due to problems with calcification, thromboembolic complications and insufficient mechanical properties of the leaflets which are the cause of malfunctioning during the stages of opening/closing. They have however found use which is limited to devices for ventricular assistance for temporary use. Use in these systems is in fact less critical because they are intended for temporary paracorporeal use (months, or at most very few years) and the patients nevertheless always receive anticoagulant therapy.
Ideally a PHV should combine the service life of mechanical valves and the haemocompatibility of biological valves, overcoming the disadvantages, mainly the thrombogenicity of mechanical valves and the poor service life of biological valves. Also new emerging therapeutic alternatives such as valve replacement by a minimally invasive percutaneous approach, which require devices capable of being collapsed and introduced within small diameter catheters, have attracted greater attention to the PHV option. Another new concept in valve replacement therapy, the tissue engineering of PHV, which uses biodegradable synthetic polymers as a scaffold, has recently increased interest in polymer materials.
Choice of material is important for the development of PHV because the material helps to provide the valve with durability and biocompatibility, so as to overcome the clinical problems associated with both mechanical and biological valves, such as thromboembolic events, undesired events due to anticoagulants and premature failure, providing improved haemodynamic functionality and service life.
Thus because a PHV is becoming a valid alternative option to valve replacement therapy the polymer selected must not only have acceptable characteristics with regard to biostability, haemocompatibility, anti-thrombogenicity, resistance to degradation and calcification, it must also have good affinity for endothelial cells.
Various synthetic polymers have been used as materials for valve leaflets, including inert synthetics such as silicone and polyolefin rubber, but these have proved to have an inadequate service life and have therefore been subsequently abandoned. Polytetrafluoroethylene (PTFE) has not had success as a material for PHV for similar reasons, in that it has given rise to a high incidence of thrombosis and calcification.
Polyurethanes (PU) are among the most popular and successful materials for biomedical applications. This class of polymer materials in fact has a number of favourable properties deriving from a two-stage microstructure consisting of rigid crystalline segments and soft elastomer segments, the ratio between which gives rise to important properties of the material such as rigidity. The rigid segments are formed by the reaction of a diisocyanate with a short chain diol or diamine (“chain extenders”) typically 1,4-butanediol or ethylene diamine. The soft segments are formed by the reaction of diisocyanate with high molecular weight polyols typically within the range of 1000-2000 Daltons, such as polyethers, polyesters or polycarbonates. Their versatile characteristics, such as for example haemocompatibility and improved haemodynamic and mechanical properties, make PU useful materials for the development of cardiovascular devices.
However the main disadvantage associated with long-term applications is their low biostability, which is mainly caused by their susceptibility to degradation. The degradation of PUs is brought about by oxidation, acid hydrolysis or enzyme sequences and results in the loss of mechanical properties and eventually the creation of lacerations or cracks in the valve leaflets. The second and more serious disadvantage of PU is their tendency to calcification, which remains an appreciable obstacle to their use in long-term implants.
In order to deal with these problems efforts have been made to improve the properties of polyurethanes by modifying the soft segments, which are considered to be the most vulnerable components. Up to now three main types of PU with different soft segments, that is polyester urethanes (PEsU), polyether urethanes (PEtU) and polycarbonate urethanes (PCU) have been developed and consequently tested in biomedical applications.
The first generation of PU used in medical devices were the PEsU, but these proved unsuitable for long-term implants because of rapid hydrolysis of the soft polyester segment. PEtU on the contrary have excellent stability to hydrolysis and have therefore replaced PEsU in implantable medical devices for a couple of decades.
However, recently various studies have demonstrated that the soft segment of polyether is also susceptible to oxidative degradation and suffers environmental stress cracking under the conditions of in vivo implants.
Subsequently the third class of PUs, PCUs, have been tested and have demonstrated that they have greater stability to oxidation. In comparison with PEtUs the degree of biodegradation of PCUs has proved to be significantly lower and restricted to a thin surface layer. Replacements of the chemical structure of PUs have also been made in an attempt to increase their biostability. The, binding of biodegradation-resistant molecules to the polymer has proved an effective method for increasing the biostability of PUs. For example attempts have been made to incorporate polydimethylsiloxane (PDMS) (a molecule which imparts good thermal and oxidative stability) into the PU chain in the presence of polyhexamethylene oxide (PHMO), which is a compatibilising polyether facilitating incorporation of the non-polar PDMS macrodiol into the PU.
The idea underlying this proposed patent is the development of a new design of PHV with a geometry similar to that of a natural aortic valve (and therefore that of biological valve prostheses) which is not subject to calcification, has a long service life and a morphology such as to reduce the thromboembolic complications due to its interaction with blood flow and with cardiac and vascular tissue to a minimum.
The valve, a single body incorporated with the supporting stent, is made using a semi-interpenetrating polymer network (semi-IPN) newly synthesised on the basis of a co-polymer of poly(carbonate-urethane) (PCU) and polymethylsiloxane (PDMS), cross-linked with a functionalised silicone (functional-PDMS) and capable of combining the best mechanical strength, biocompatibility and superior biostability properties of PCU with the excellent haemocompatibility and calcification-resistance properties of silicone (PDMS) as a material of manufacture.
The presence of silicone in the polyurethane chain, together with that of the cross-linking silicone forming the semi-IPN makes it possible to vary the flexibility (flex-life) and biodegradation resistance of the new valves.
As far as the design of PHV is concerned, it is well known that the structural anatomy of the natural valve plays an essential part in its operating function, providing a suitable and stable structure with specific anatomical and histological characteristics. In view of the complex anatomy of natural valves it is difficult to create structures which have the precise anatomical and functional characteristics of a native valve. However, unlike their biological counterparts, valves with synthetic leaflets can be designed in virtually any form, and this emphasises the possible importance of structural design strategies.
The process of valve manufacture is also an essential factor influencing the performance of PHV, as its effect on the durability of valves and their haemodynamic functions has been demonstrated.
Different methods of producing PHV have been investigated, such as deep coating, film-fabrication, cavity moulding and injection moulding.
Deep-coating implies the use of a specifically designed former which undergoes repeated cycles of immersion in the polymer solution and subsequent consolidation in air or in a dry air stove until the desired thickness is achieved. The concentration of the polymer solution may vary according to the polymer chosen and the stage of manufacture. Normally the deep-coating process comprises repeated immersion in a low concentration polymer solution.
The great disadvantage of this method is that it is difficult to control the thickness distribution in the leaflet precisely. Some have proposed that leaflets should be made using a single immersion in concentrated polymer solution. This would allow more accurate reproducibility and would reduce dependence on the operator to a minimum, but because of the concentrated polymer solution undesired densification of the material in different portions of the leaflets could occur.
In film-fabrication, polymer films are deposited up to a particular thickness and the leaflets are produced by cutting the film to the desired shape. The leaflets are then anchored to the valve support structure, which is manufactured separately, through the use of solvents. Finally a thermal shaping process is used to obtain the desired valve geometry. A potential disadvantage of this technique lies in the weaknesses which may be created at the point where the leaflets are anchored to the structure of the valve because of the use of polymer solvent.
Cavity moulding uses a cavity mould comprising a static portion (female) and a moving portion (male); the mould is used to manufacture the entire valve structure through introducing hot polymer, after which the sealed mould is placed in a water bath and subjected to alternating freeze/thaw cycles to form a thin polymer film.
The disadvantage of this method lies in the fact that the material has to be subjected to different thermal cycles which could affect the fatigue resistance characteristics of the valve in a manner which is difficult to foresee.
In injection moulding an injection moulding machine is used to manufacture the valve leaflets in a partly open position on a former, after which repeated baths of hot and cold water are applied in order to produce the final valve. Here again the repeated thermal cycles may affect the mechanical and biostability characteristics of the valve.
Understandably these limitations make it difficult to manufacture heart valves having high durability and haemocompatibility standards.